Label-free electronic probing of nucleic acids and proteins at the nanoscale using the nanoneedle biosensor Rahim Esfandyarpour, Mehdi Javanmard, Zahra Koochak, Hesaam Esfandyarpour, James S. Harris, and Ronald W. Davis Citation: Biomicrofluidics 7, 044114 (2013); doi: 10.1063/1.4817771 View online: http://dx.doi.org/10.1063/1.4817771 View Table of Contents: http://scitation.aip.org/content/aip/journal/bmf/7/4?ver=pdfcov Published by the AIP Publishing Articles you may be interested in High sensitivity and high Q-factor nanoslotted parallel quadrabeam photonic crystal cavity for real-time and labelfree sensing Appl. Phys. Lett. 105, 063118 (2014); 10.1063/1.4867254 Erratum: “Label-free electronic probing of nucleic acids and proteins at the nanoscale using the nanoneedle biosensor” [Biomicrofluidics 7, 044114 (2013)] Biomicrofluidics 8, 029901 (2014); 10.1063/1.4869375 Publisher's Note: “Label-free electronic probing of nucleic acids and proteins at the nanoscale using the nanoneedle biosensor” [Biomicrofluidics 7, 044114 (2013)] Biomicrofluidics 7, 049901 (2013); 10.1063/1.4819277 A new type of optical biosensor from DNA wrapped semiconductor graphene ribbons J. Appl. Phys. 111, 114703 (2012); 10.1063/1.4728196 Self-assembled nanotube field-effect transistors for label-free protein biosensors J. Appl. Phys. 104, 074310 (2008); 10.1063/1.2988274
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BIOMICROFLUIDICS 7, 044114 (2013)
Label-free electronic probing of nucleic acids and proteins at the nanoscale using the nanoneedle biosensor Rahim Esfandyarpour,1,2,a) Mehdi Javanmard,2 Zahra Koochak,3 Hesaam Esfandyarpour,1 James S. Harris,1 and Ronald W. Davis2 1
Center for Integrated Systems, Department of Electrical Engineering, Stanford University, 855 California Ave., Palo Alto, California 94304, USA 2 Stanford Genome Technology Center, 855 California Ave., Palo Alto, California 94304, USA 3 University of California Santa Cruz, Santa Cruz, California 95064, USA (Received 10 June 2013; accepted 24 July 2013; published online 6 August 2013; publisher error corrected 15 August 2013)
Detection of proteins and nucleic acids is dominantly performed using optical fluorescence based techniques, which are more costly and timely than electrical detection due to the need for expensive and bulky optical equipment and the process of fluorescent tagging. In this paper, we discuss our study of the electrical properties of nucleic acids and proteins at the nanoscale using a nanoelectronic probe we have developed, which we refer to as the Nanoneedle biosensor. The nanoneedle consists of four thin film layers: a conductive layer at the bottom acting as an electrode, an oxide layer on top, and another conductive layer on top of that, with a protective oxide above. The presence of proteins and nucleic acids near the tip results in a decrease in impedance across the sensing electrodes. There are three basic mechanisms behind the electrical response of DNA and protein molecules in solution under an applied alternating electrical field. The first change stems from modulation of the relative permittivity at the interface. The second mechanism is the formation and relaxation of the induced dipole moment. The third mechanism is the tunneling of electrons through the biomolecules. The results presented in this paper can be extended to develop low cost point-of-care diagnostic assays for the C 2013 AIP Publishing LLC. [http://dx.doi.org/10.1063/1.4817771] clinical setting. V
I. INTRODUCTION
Direct electrical detection of biomolecules without the need for any labeling can help greatly advance point-of-care diagnostics. Applications include the study of virology,1–3 ligand fishing,4,5 bacteriology,6–8 apoptosis,9 cell biology and adhesion,10,11 epitope mapping,12–14 signal transduction,15,16 immune regulation,17 nucleic acid–nucleic acid interactions,18–20 and nucleic acid–protein protein interactions,21,22 and study of post-translational modifications.23,24 Detection of proteins and nucleic acids is often performed using optical fluorescence based techniques, which are more costly and timely than electrical detection due to the need for expensive and bulky optical equipment and the process of fluorescent tagging. Thus, a robust label-free electrical detection technique can provide for a promising solution in lowering both reagent costs and instrumentation costs. Optical detection of nucleic acids in miniaturized systems is also challenging because the signal originates from dye molecules in solution, and thus the strength of the signal scales with sample volume. Therefore, there is a direct conflict between the goals of obtaining a strong optical signal and reducing reagent consumption in a microfluidic system. Furthermore, optical readout requires that PCR product markers such as Sybr Green and Taqman probes be added to the reagents. This process of labeling and adding reagents makes
a)
Email: [email protected]. Telephone: þ1-650-387-5976.
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7, 044114-1
C 2013 AIP Publishing LLC V
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Biomicrofluidics 7, 044114 (2013)
this process of real time PCR unsuitable as a point-of-use diagnostic technique in the clinical setting. Protein detection is typically performed using the sandwich ELISA technique, which involves several steps of incubating test sample, then a polyclonal antibody, and then finally a secondary antibody tagged with a fluorescent or luminescent label, with several wash steps in between. A label-free technique, which could directly detect the binding of a target protein to the surface antibody would be much more suitable as a point of care diagnostic. Various labelfree nanoelectronic25–28 sensors including nanowires have been demonstrated exhibiting femtomolar detection limits.29 The detection limit or the minimum detectable concentration of target biomarkers in the test sample is dependent on two parameters: the transducer sensitivity (the minimum number of binding events on the sensor surface required to generate sensor response greater than the noise level) and the capture rate of target molecules on the surface of the sensor. Both of these parameters are affected by the flow rate, diffusion time of target molecules, and the sensor geometry. In general, most of the electrical impedance based biosensors suffers from low transducer signal to noise ratio due to various noise processes in the system such as flicker (1/f), JohnsonNyquist noise, and also the noise resulting from the amplification circuitry. The contribution of these noise sources are relatively higher at frequencies below 100 Hz where electrical impedance measurements are typically made, since the desired signal to be measured results from modulation of the double layer as target biomarkers bind to the sensor surface. To overcome these various problems mentioned, we propose a novel method of directly measuring the electrical response of the DNA and protein molecules of interest at the nanoscale by using a novel ultra-sensitive, real time, label-free sensing platform, which we refer to as the nanoneedle biosensor. The nanoneedle biosensor structure consists of three thin-film layers, as shown in Fig. 1. There are two conductive layers with an insulator layer in between. The interface of this middle oxide layer with the electrolyte is the active region of the sensor. A protective oxide layer is deposited above the topmost electrode. Underneath the bottom electrode, there is an oxide layer, which is thermally grown insulating the first electrode from the substrate. One of the advantages of this sensor is the ability to directly measure biomolecular binding as a function of time (real-time). The presence of biomolecules in the active sensing region of the needle tip results in modulation of the measured impedance in real time. This can be useful for measuring kinetic constants for various biomolecular species. Also, transducer sensitivity is
FIG. 1. (a) Schematic of nanoneedle biosensor three-dimensional and side view of horizontal nanoneedles (Not to Scale). (b) Optical micrograph of bird’s eye view of aluminum-polysiliconhybridnanoneedle biosensor. (c) SEM image of the tip of a nanoneedles biosensor; 1 & 3 are the electrodes; 2 is the oxide in between the electrodes
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improved with the nanoneedle biosensor compared to standard micro-electrode label-free impedance sensors due to miniaturization of the nanoneedle tip, resulting in high sensitivity for detection of small numbers of molecules. The sensing area of this sensor is a nano-sized area, which is in the size range of biomolecules of interest. As a result, a small number of binding events in the active sensing area is sufficient to modulate the impedance at the sensor tip to a level greater than the noise resulting in high sensitivity. In addition, the suspended geometry of a nanoneedle in a micro-channel results in diffusion taking place in three dimensions unlike most electrical biosensors which have a planar structure(two dimensional) thus diffusion taking place in only two dimensions. This results in a higher hit rate of target molecules to the probe molecules on the sensor surface and thus a faster detection platform. Another advantage of the nanoneedles biosensor, due to its rigid yet high aspect ratio solidstate structure, is the ability to measure protein and nucleic acid levels directly in-vivo inside a living cell. A thin functionalized needle can be inserted into a living cell, and impedance measurements can be made directly as proteins bind to the needle surface. This can be used for many different applications such as measuring protein expression for the purposes of drug screening. In addition to the above-mentioned advantages, the ability to fabricate an array of needles onto a substrate potentially enables high throughput sensing. Fabrication of an array of nanoneedles with the state-of-the-art nanofabrication techniques makes it possible to monitor various binding events simultaneously in over a large area. On chip integration of the sensors with CMOS amplification electronics can further improve the signal to noise ratio of the sensor resulting in a single portable device suitable for point-of-care diagnostics. II. DEVICE DESIGN FACTORS
Various thicknesses and geometrical designs have been fabricated and tested. The sensor design used in this study consists of electrodes 100 nm thick and a middle oxide layer 30 nm thick. The top protective oxide layer thickness is 20 nm and the bottom oxide layer thickness is 250 nm. The width of the nanoneedle tip is 5 lm. For specificity, probe molecules (e.g., DNA molecule or protein) can be immobilized on the tip of the nanoneedles. The binding of target molecules to the probe molecules modulates the impedance between the electrodes. In order to maximize the effect of the impedance at the interface of the sensor compared to the total measured impedance, it was necessary to reduce the parasitic impedance resulting from the resistance of the trace and also the body capacitance of the traces leading from the bonding pad to the sensor. To achieve this goal, the sensor has an aluminum/polysilicon hybrid structure, where the sensor electrode is polysilicon, and the trace leading up to it is aluminum, which has a higher conductivity than polysilicon and can minimize the trace resistance leading up to the sensor. And also the pads are separated out from each other to eliminate the body capacitance between the traces. However, the sensing area will still be polysilicon in order to ensure compatibility with surface chemistries optimized for high immobilization capacity of probe molecules. III. RESULTS AND DISCUSSIONS
In order to demonstrate the utility of the nanoneedle for label-free sensing while maintaining a high signal-noise ratio we studied the electrical response of the sensor for various types of biological agents. We tested the electrical response of two types of biomolecules: nucleic acids and proteins. In this manuscript, first we present the results of the experimental characterization of the device, then we discuss the theory behind the response and discuss the various physical mechanisms involved, and afterwards we demonstrate the utility of the sensor for affinity based protein sensing. A. Nucleic acids
We initially studied the electrical response of nucleic acids using the nanoneedle sensor. As observed in Fig. 2(a), the presence of single stranded DNA (20 base pairs long) modulates the measured impedance. Various concentrations of DNA were injected onto our sensor surface
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FIG. 2. Presence of single stranded DNA modulates the measured impedance. (a) Various concentrations of DNA were injected onto our sensor surface sequentially. Between every step that DNA was added we dried out the measurement well. As the concentration of DNA in the solution decreases, the measured impedance increases getting closer and closer to the baseline value. (b) Impedance change plotted with error bars over three measurements per point.
sequentially. Between every step that DNA was added we aspirated the measurement well. The DNA is in free solution and unlikely to adsorb to the surface. As the concentration of DNA in the solution decreases, the measured impedance increases getting closer and closer to the baseline value. Fig. 2(b) shows the measurement performed with error bars. We acknowledge, however, that during the aspiration step it is possible that residues of DNA remain on the substrate resulting in the actual concentration being somewhat higher than the concentration we injected. For example, for the final measured concentration (0.6 fM which equates to 1800 DNA molecules in 5 ll of volume), it is possible that the actual number of molecules is somewhat higher. This observation of increase in conductivity with higher DNA concentration was counter intuitive and the results are contrary to the observations traditionally made with electrical impedance biosensors where the presence of DNA generally results in an increase in impedance as a result of modulation of the double layer.30 We will discuss the mechanisms affecting the modulation of the impedance level in Sec. IV; however, in order to confirm the behavior of the sensor even further, we also performed a next set of experiments with proteins. B. Proteins
For protein experiments, we measured the electrical response of unconjugated streptavidin in free solution (Fig. 3). Since streptavidin has an isoelectric point of 4, and silicon oxide has an isoelectric point of 3, both will have a negatively charged surface in water, thus minimizing adsorption of molecules to the surface. We measured the impedance across the electrodes as varying concentrations of streptavidin was injected onto the sensor surface. Fig. 3(a) shows representative results of this experiment, which was performed three times. The behavior of streptavidin is similar to the electrical response of the DNA. Higher concentration of protein resulted in lower levels of impedance between the electrodes. Again, similar to our results with DNA, this goes contrary to the behavior of traditional micro-electrode sensors where the presence of biomolecules at the surface results in an effective decrease in the double layer capacitance, thus resulting in an increase in impedance.30 C. Polystyrene beads
As a control experiment to verify the physical mechanism resulting in the electrical response of the sensor we injected polystyrene beads. Polystyrene beads have fully insulative electrical properties, and also a lower dielectric constant (2.6) compared to water (80). As shown in Fig. 3(b)) an increase in impedance was observed as opposed to the decrease in impedance seen with injection of proteins and DNA. Fig. 3(c) shows the impedance measurements repeated several times with error bars included. The presence of the beads on the sensor surface results in both an increase in resistance across the electrodes, and also a decrease in the sensor
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FIG. 3. Presence of non-adsorbed streptavidin protein modulates the measured impedance. (a) Various concentrations of streptavidin were injected onto our sensor surface sequentially: (1) water, (2) 250 ng/ml streptavidin, (3) 25 ug/ml streptavidin, (4) 25 mg/ml streptavidin. Between every step that protein was added we dried out the measurement well. As the concentration of protein in the solution decreases, the measured impedance increases getting closer and closer to the baseline value. (b) Impedance response of polystyrene beads injected onto sensor resulting in increase in impedance. (c) Impedance change plotted with error bars over three measurements per point.
surface capacitances, all of which contribute to an increase in impedance. This behavior is in line with that of traditional micro-scale electrodes. This implies that the charge and the relative dielectric constant of the biomolecules play an important role in the behavior observed for nucleic acids and proteins. To reconcile this contradiction of our protein/nucleic acid results with the traditional micro-sensors, we theoretically and experimentally characterized the various parasitic components of our circuit model at the sensor-electrolyte interface in our system. IV. MODELING
To understand the results, we developed a full circuit model to characterize the interface of the nanoneedle sensor with the electrolyte as shown in Fig. 4(a). In the full model, our system contains several parasitic impedance components. From here on, we are referring to the parasitic impedances at the electrode-electrolyte interface and the electrolyte itself (rather than the components of the body of the sensor). Above the electrodes and the insulator, we have a double layer resulting from accumulation of ions in the electrolyte at each surface (metal and oxide). The electrical double layer consists of two layers, the stern layer, which is an adsorbed fixed layer, and also a diffuse layer. The stern layer consists of the ions that are adsorbed to the surface and are estimated to be concentrated roughly 1 nm from the sensor surface. The diffuse layer results because of the condition of charge neutrality. That is, another layer of ions accumulates in order to neutralize the charge in the stern layer. The thickness of the stern layer is generally concentration independent, whereas that of the diffuse layer is related inversely to the electrolyte concentration. The stern layer and diffuse layer can be represented as two capacitors Cads and Cdiff
Cdif f
rffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi Az2 e2 ee0 Ci NA ; ¼ ee0 K ¼ kT
(1)
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FIG. 4. Circuit model of the nanoneedle sensor-electrolyte interface. (a) Full model where CW represents the fringing capacitance at the sensor interface. Rw represents the resistance across the double layer (on top of the insulator) between the electrodes. Cdl represents the double layer capacitance on each electrode surface. Rf represents the tunneling resistance or the electron transfer resistance from the electrode into the bulk solution. Rb represents the bulk resistance of the electrolyte. Rtr represents the trace resistance of the electrode leading up to the bonding pads. Cb represents the body capacitance between the electrodes along the body of the sensor. (b) Simplified model which is valid at f ¼ 15 kHz.
where z is the ion valence number, A is a constant, Ci is the concentration of the ionic solution (mol/liter), e the charge of an electron (1.6 1019), NA is Avogadro’s number, ee0 is the absolute dielectric constant of the sample, k is the Boltzmann constant, and T is the absolute temperature. The Cdl is the result of both Cads and Cdiff which are connected in series 1 1 1 ¼ þ : Cdl Cads CDif f
(2)
At high salt concentrations (>10 mM), the diffuse layer size becomes comparable to that of the stern layer. However, when the salt concentration is low (
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BIOMICROFLUIDICS 7, 044114 (2013)
Label-free electronic probing of nucleic acids and proteins at the nanoscale using the nanoneedle biosensor Rahim Esfandyarpour,1,2,a) Mehdi Javanmard,2 Zahra Koochak,3 Hesaam Esfandyarpour,1 James S. Harris,1 and Ronald W. Davis2 1
Center for Integrated Systems, Department of Electrical Engineering, Stanford University, 855 California Ave., Palo Alto, California 94304, USA 2 Stanford Genome Technology Center, 855 California Ave., Palo Alto, California 94304, USA 3 University of California Santa Cruz, Santa Cruz, California 95064, USA (Received 10 June 2013; accepted 24 July 2013; published online 6 August 2013; publisher error corrected 15 August 2013)
Detection of proteins and nucleic acids is dominantly performed using optical fluorescence based techniques, which are more costly and timely than electrical detection due to the need for expensive and bulky optical equipment and the process of fluorescent tagging. In this paper, we discuss our study of the electrical properties of nucleic acids and proteins at the nanoscale using a nanoelectronic probe we have developed, which we refer to as the Nanoneedle biosensor. The nanoneedle consists of four thin film layers: a conductive layer at the bottom acting as an electrode, an oxide layer on top, and another conductive layer on top of that, with a protective oxide above. The presence of proteins and nucleic acids near the tip results in a decrease in impedance across the sensing electrodes. There are three basic mechanisms behind the electrical response of DNA and protein molecules in solution under an applied alternating electrical field. The first change stems from modulation of the relative permittivity at the interface. The second mechanism is the formation and relaxation of the induced dipole moment. The third mechanism is the tunneling of electrons through the biomolecules. The results presented in this paper can be extended to develop low cost point-of-care diagnostic assays for the C 2013 AIP Publishing LLC. [http://dx.doi.org/10.1063/1.4817771] clinical setting. V
I. INTRODUCTION
Direct electrical detection of biomolecules without the need for any labeling can help greatly advance point-of-care diagnostics. Applications include the study of virology,1–3 ligand fishing,4,5 bacteriology,6–8 apoptosis,9 cell biology and adhesion,10,11 epitope mapping,12–14 signal transduction,15,16 immune regulation,17 nucleic acid–nucleic acid interactions,18–20 and nucleic acid–protein protein interactions,21,22 and study of post-translational modifications.23,24 Detection of proteins and nucleic acids is often performed using optical fluorescence based techniques, which are more costly and timely than electrical detection due to the need for expensive and bulky optical equipment and the process of fluorescent tagging. Thus, a robust label-free electrical detection technique can provide for a promising solution in lowering both reagent costs and instrumentation costs. Optical detection of nucleic acids in miniaturized systems is also challenging because the signal originates from dye molecules in solution, and thus the strength of the signal scales with sample volume. Therefore, there is a direct conflict between the goals of obtaining a strong optical signal and reducing reagent consumption in a microfluidic system. Furthermore, optical readout requires that PCR product markers such as Sybr Green and Taqman probes be added to the reagents. This process of labeling and adding reagents makes
a)
Email: [email protected]. Telephone: þ1-650-387-5976.
1932-1058/2013/7(4)/044114/13/$30.00
7, 044114-1
C 2013 AIP Publishing LLC V
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044114-2
Esfandyarpour et al.
Biomicrofluidics 7, 044114 (2013)
this process of real time PCR unsuitable as a point-of-use diagnostic technique in the clinical setting. Protein detection is typically performed using the sandwich ELISA technique, which involves several steps of incubating test sample, then a polyclonal antibody, and then finally a secondary antibody tagged with a fluorescent or luminescent label, with several wash steps in between. A label-free technique, which could directly detect the binding of a target protein to the surface antibody would be much more suitable as a point of care diagnostic. Various labelfree nanoelectronic25–28 sensors including nanowires have been demonstrated exhibiting femtomolar detection limits.29 The detection limit or the minimum detectable concentration of target biomarkers in the test sample is dependent on two parameters: the transducer sensitivity (the minimum number of binding events on the sensor surface required to generate sensor response greater than the noise level) and the capture rate of target molecules on the surface of the sensor. Both of these parameters are affected by the flow rate, diffusion time of target molecules, and the sensor geometry. In general, most of the electrical impedance based biosensors suffers from low transducer signal to noise ratio due to various noise processes in the system such as flicker (1/f), JohnsonNyquist noise, and also the noise resulting from the amplification circuitry. The contribution of these noise sources are relatively higher at frequencies below 100 Hz where electrical impedance measurements are typically made, since the desired signal to be measured results from modulation of the double layer as target biomarkers bind to the sensor surface. To overcome these various problems mentioned, we propose a novel method of directly measuring the electrical response of the DNA and protein molecules of interest at the nanoscale by using a novel ultra-sensitive, real time, label-free sensing platform, which we refer to as the nanoneedle biosensor. The nanoneedle biosensor structure consists of three thin-film layers, as shown in Fig. 1. There are two conductive layers with an insulator layer in between. The interface of this middle oxide layer with the electrolyte is the active region of the sensor. A protective oxide layer is deposited above the topmost electrode. Underneath the bottom electrode, there is an oxide layer, which is thermally grown insulating the first electrode from the substrate. One of the advantages of this sensor is the ability to directly measure biomolecular binding as a function of time (real-time). The presence of biomolecules in the active sensing region of the needle tip results in modulation of the measured impedance in real time. This can be useful for measuring kinetic constants for various biomolecular species. Also, transducer sensitivity is
FIG. 1. (a) Schematic of nanoneedle biosensor three-dimensional and side view of horizontal nanoneedles (Not to Scale). (b) Optical micrograph of bird’s eye view of aluminum-polysiliconhybridnanoneedle biosensor. (c) SEM image of the tip of a nanoneedles biosensor; 1 & 3 are the electrodes; 2 is the oxide in between the electrodes
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044114-3
Esfandyarpour et al.
Biomicrofluidics 7, 044114 (2013)
improved with the nanoneedle biosensor compared to standard micro-electrode label-free impedance sensors due to miniaturization of the nanoneedle tip, resulting in high sensitivity for detection of small numbers of molecules. The sensing area of this sensor is a nano-sized area, which is in the size range of biomolecules of interest. As a result, a small number of binding events in the active sensing area is sufficient to modulate the impedance at the sensor tip to a level greater than the noise resulting in high sensitivity. In addition, the suspended geometry of a nanoneedle in a micro-channel results in diffusion taking place in three dimensions unlike most electrical biosensors which have a planar structure(two dimensional) thus diffusion taking place in only two dimensions. This results in a higher hit rate of target molecules to the probe molecules on the sensor surface and thus a faster detection platform. Another advantage of the nanoneedles biosensor, due to its rigid yet high aspect ratio solidstate structure, is the ability to measure protein and nucleic acid levels directly in-vivo inside a living cell. A thin functionalized needle can be inserted into a living cell, and impedance measurements can be made directly as proteins bind to the needle surface. This can be used for many different applications such as measuring protein expression for the purposes of drug screening. In addition to the above-mentioned advantages, the ability to fabricate an array of needles onto a substrate potentially enables high throughput sensing. Fabrication of an array of nanoneedles with the state-of-the-art nanofabrication techniques makes it possible to monitor various binding events simultaneously in over a large area. On chip integration of the sensors with CMOS amplification electronics can further improve the signal to noise ratio of the sensor resulting in a single portable device suitable for point-of-care diagnostics. II. DEVICE DESIGN FACTORS
Various thicknesses and geometrical designs have been fabricated and tested. The sensor design used in this study consists of electrodes 100 nm thick and a middle oxide layer 30 nm thick. The top protective oxide layer thickness is 20 nm and the bottom oxide layer thickness is 250 nm. The width of the nanoneedle tip is 5 lm. For specificity, probe molecules (e.g., DNA molecule or protein) can be immobilized on the tip of the nanoneedles. The binding of target molecules to the probe molecules modulates the impedance between the electrodes. In order to maximize the effect of the impedance at the interface of the sensor compared to the total measured impedance, it was necessary to reduce the parasitic impedance resulting from the resistance of the trace and also the body capacitance of the traces leading from the bonding pad to the sensor. To achieve this goal, the sensor has an aluminum/polysilicon hybrid structure, where the sensor electrode is polysilicon, and the trace leading up to it is aluminum, which has a higher conductivity than polysilicon and can minimize the trace resistance leading up to the sensor. And also the pads are separated out from each other to eliminate the body capacitance between the traces. However, the sensing area will still be polysilicon in order to ensure compatibility with surface chemistries optimized for high immobilization capacity of probe molecules. III. RESULTS AND DISCUSSIONS
In order to demonstrate the utility of the nanoneedle for label-free sensing while maintaining a high signal-noise ratio we studied the electrical response of the sensor for various types of biological agents. We tested the electrical response of two types of biomolecules: nucleic acids and proteins. In this manuscript, first we present the results of the experimental characterization of the device, then we discuss the theory behind the response and discuss the various physical mechanisms involved, and afterwards we demonstrate the utility of the sensor for affinity based protein sensing. A. Nucleic acids
We initially studied the electrical response of nucleic acids using the nanoneedle sensor. As observed in Fig. 2(a), the presence of single stranded DNA (20 base pairs long) modulates the measured impedance. Various concentrations of DNA were injected onto our sensor surface
This article is copyrighted as indicated in the article. Reuse of AIP content is subject to the terms at: http://scitation.aip.org/termsconditions. Downloaded to IP: 129.49.23.145 On: Thu, 18 Dec 2014 06:26:36
044114-4
Esfandyarpour et al.
Biomicrofluidics 7, 044114 (2013)
FIG. 2. Presence of single stranded DNA modulates the measured impedance. (a) Various concentrations of DNA were injected onto our sensor surface sequentially. Between every step that DNA was added we dried out the measurement well. As the concentration of DNA in the solution decreases, the measured impedance increases getting closer and closer to the baseline value. (b) Impedance change plotted with error bars over three measurements per point.
sequentially. Between every step that DNA was added we aspirated the measurement well. The DNA is in free solution and unlikely to adsorb to the surface. As the concentration of DNA in the solution decreases, the measured impedance increases getting closer and closer to the baseline value. Fig. 2(b) shows the measurement performed with error bars. We acknowledge, however, that during the aspiration step it is possible that residues of DNA remain on the substrate resulting in the actual concentration being somewhat higher than the concentration we injected. For example, for the final measured concentration (0.6 fM which equates to 1800 DNA molecules in 5 ll of volume), it is possible that the actual number of molecules is somewhat higher. This observation of increase in conductivity with higher DNA concentration was counter intuitive and the results are contrary to the observations traditionally made with electrical impedance biosensors where the presence of DNA generally results in an increase in impedance as a result of modulation of the double layer.30 We will discuss the mechanisms affecting the modulation of the impedance level in Sec. IV; however, in order to confirm the behavior of the sensor even further, we also performed a next set of experiments with proteins. B. Proteins
For protein experiments, we measured the electrical response of unconjugated streptavidin in free solution (Fig. 3). Since streptavidin has an isoelectric point of 4, and silicon oxide has an isoelectric point of 3, both will have a negatively charged surface in water, thus minimizing adsorption of molecules to the surface. We measured the impedance across the electrodes as varying concentrations of streptavidin was injected onto the sensor surface. Fig. 3(a) shows representative results of this experiment, which was performed three times. The behavior of streptavidin is similar to the electrical response of the DNA. Higher concentration of protein resulted in lower levels of impedance between the electrodes. Again, similar to our results with DNA, this goes contrary to the behavior of traditional micro-electrode sensors where the presence of biomolecules at the surface results in an effective decrease in the double layer capacitance, thus resulting in an increase in impedance.30 C. Polystyrene beads
As a control experiment to verify the physical mechanism resulting in the electrical response of the sensor we injected polystyrene beads. Polystyrene beads have fully insulative electrical properties, and also a lower dielectric constant (2.6) compared to water (80). As shown in Fig. 3(b)) an increase in impedance was observed as opposed to the decrease in impedance seen with injection of proteins and DNA. Fig. 3(c) shows the impedance measurements repeated several times with error bars included. The presence of the beads on the sensor surface results in both an increase in resistance across the electrodes, and also a decrease in the sensor
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044114-5
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Biomicrofluidics 7, 044114 (2013)
FIG. 3. Presence of non-adsorbed streptavidin protein modulates the measured impedance. (a) Various concentrations of streptavidin were injected onto our sensor surface sequentially: (1) water, (2) 250 ng/ml streptavidin, (3) 25 ug/ml streptavidin, (4) 25 mg/ml streptavidin. Between every step that protein was added we dried out the measurement well. As the concentration of protein in the solution decreases, the measured impedance increases getting closer and closer to the baseline value. (b) Impedance response of polystyrene beads injected onto sensor resulting in increase in impedance. (c) Impedance change plotted with error bars over three measurements per point.
surface capacitances, all of which contribute to an increase in impedance. This behavior is in line with that of traditional micro-scale electrodes. This implies that the charge and the relative dielectric constant of the biomolecules play an important role in the behavior observed for nucleic acids and proteins. To reconcile this contradiction of our protein/nucleic acid results with the traditional micro-sensors, we theoretically and experimentally characterized the various parasitic components of our circuit model at the sensor-electrolyte interface in our system. IV. MODELING
To understand the results, we developed a full circuit model to characterize the interface of the nanoneedle sensor with the electrolyte as shown in Fig. 4(a). In the full model, our system contains several parasitic impedance components. From here on, we are referring to the parasitic impedances at the electrode-electrolyte interface and the electrolyte itself (rather than the components of the body of the sensor). Above the electrodes and the insulator, we have a double layer resulting from accumulation of ions in the electrolyte at each surface (metal and oxide). The electrical double layer consists of two layers, the stern layer, which is an adsorbed fixed layer, and also a diffuse layer. The stern layer consists of the ions that are adsorbed to the surface and are estimated to be concentrated roughly 1 nm from the sensor surface. The diffuse layer results because of the condition of charge neutrality. That is, another layer of ions accumulates in order to neutralize the charge in the stern layer. The thickness of the stern layer is generally concentration independent, whereas that of the diffuse layer is related inversely to the electrolyte concentration. The stern layer and diffuse layer can be represented as two capacitors Cads and Cdiff
Cdif f
rffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffiffi Az2 e2 ee0 Ci NA ; ¼ ee0 K ¼ kT
(1)
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044114-6
Esfandyarpour et al.
Biomicrofluidics 7, 044114 (2013)
FIG. 4. Circuit model of the nanoneedle sensor-electrolyte interface. (a) Full model where CW represents the fringing capacitance at the sensor interface. Rw represents the resistance across the double layer (on top of the insulator) between the electrodes. Cdl represents the double layer capacitance on each electrode surface. Rf represents the tunneling resistance or the electron transfer resistance from the electrode into the bulk solution. Rb represents the bulk resistance of the electrolyte. Rtr represents the trace resistance of the electrode leading up to the bonding pads. Cb represents the body capacitance between the electrodes along the body of the sensor. (b) Simplified model which is valid at f ¼ 15 kHz.
where z is the ion valence number, A is a constant, Ci is the concentration of the ionic solution (mol/liter), e the charge of an electron (1.6 1019), NA is Avogadro’s number, ee0 is the absolute dielectric constant of the sample, k is the Boltzmann constant, and T is the absolute temperature. The Cdl is the result of both Cads and Cdiff which are connected in series 1 1 1 ¼ þ : Cdl Cads CDif f
(2)
At high salt concentrations (>10 mM), the diffuse layer size becomes comparable to that of the stern layer. However, when the salt concentration is low (
